Heparin coatings

ABSTRACT

The invention includes a medical hydrogel made from polymerized polysaccharide macromers. The macromers are preferably polysaccharides decorated with polymerizable groups, for example, methacrylates. The macromers may also be made into polymers of at least two macromers polymerized together. These polymers are preferably multi-armed or high-molecular weight and used for medical uses, for example, making coatings on medical devices. Macromers of N-vinylpyrrolidone are also disclosed herein.

RELATED APPLICATIONS

The present application is a continuation of Ser. No. 10/179,453, filedJun. 25, 2002, which claims priority to U.S. provisional patentapplication Ser. No. 60/301,176, entitled “Polysaccharide Biomaterialsand Methods of Use Thereof”, filed Jun. 26, 2001, which are herebyincorporated herein by reference.

FIELD OF INVENTION

The invention relates to body treating compositions, especially thoseformed as a matrix. More specifically, the invention includes hydrogelmatrices of polymerized polysaccharides, especially heparin, so as tocreate a hydrogel, especially for vascular graft applications.

BACKGROUND OF THE INVENTION

Many synthetic materials have been medically used in the body, includingpolyester (e.g., DACRON™), polyethylene (e.g., milk jugs), andfluorocarbons (e.g., TEFLON™), and metals. A patient's body responds bytreating a synthetic material as an invader, although it responds onlymildly in some medical applications; for example, a metal hip implant isgenerally well tolerated. One common response to an implant is called aforeign body response in which the body forms a capsule of cells aroundthe material; the body's response to a splinter is a foreign bodyresponse. When synthetic material is used as an artificial blood vessel,for example, the blood that flows through the artificial blood vesselreacts with the synthetic material. The reaction can cause clots to formthat flow downstream that may eventually become stuck in a smallervessel; if this happens in the brain, it is called a stroke. The bloodclot can also grow on the inside of the vessel and block or severelyrestrict the blood flow. Blood's clotting mechanism is highly reactiveand, despite years of medical research, no implantable blood-contactingsynthetic material has yet been found that does not cause blood toreact.

Tubes made of polyester or fluorocarbons are currently used as largediameter blood vessels. The blood clots onto the interior walls andreduces the inside diatmeter of the vessel but the blood flow is notunduly reduced. The blood clot on the tubular wall serves as aprotective layer that elicits very little reaction from blood flowingthrough the vessel. This approach, however, does not work for smalldiameter blood vessels because the small diameter tubes are blocked whenthe blood clots onto the walls.

There are no currently known materials and techniques for manufacturingsmall diameter vascular grafts made of synthetic materials.Unfortunately, there is a great need for such grafts. One example is thecondition called deep vein thrombosis wherein a patient's veins becomeblocked. The blood drains poorly from the leg and amputation can result.Unlike heart bypass surgeries where the patient often has some bloodvessels that can be harvested from elsewhere in the body and sewn intoplace, there are few choices for replacing the long veins of the leg.

One area of synthetic biomaterials research has focused on hydrogels,materials that have a high water content and are soft and slippery. Softcontact lenses are examples of hydrogels. The materials used to make apure hydrogel might all dissolve in water but the hydrogel itself doesnot dissolve in water because the materials are cross-linked; in otherwords, the individual molecular chains are linked together like thestrands in a net or a spider's web. Hydrogels tend to elicit a milderforeign body response than other synthetic materials. Some hydrogelbiomaterials that are currently considered to be commercially useful aremade from polyethylene glycol (PEG), hyaluronic acid, and alginates.Although hydrogels tend to elicit less blood clotting than othersynthetic materials, hydrogels have not previously been successfullyused to make a small diameter vascular graft.

Scientists have also tried to use heparin to coat the inside of vasculargrafts made of synthetic materials. Heparin is a molecule that belongsto a group of molecules called polysaccharides that are polymers madefrom combinations of sugar monomers. There are many sugars; glucose andsucrose (table sugar) are two examples. Polysaccharides arenaturally-occurring polymers. Polymers are molecules built up by therepetition of smaller units that are sometimes called monomers. Polymersare typically made by special chemical schemes that make the monomerschemically react with each other to form molecular chains that can rangein length from short to very long molecules. Polymers can be assembledinto larger materials; for example, many polymers may be linked togetherto form a hydrogel.

Heparin is a polysaccharide polymer with an important property: itinterferes with key molecules in the blood clotting mechanism such thatthe blood will not clot. Coating the inside of a synthetic material tubewith heparin tends to increase the amount of time that the tube remainsopen to blood flow but, to date, small diameter vascular grafts coatedwith heparin have failed to resist blockage by blood clots for amedically useful length of time.

Heparin has been applied to materials in many ways. General strategiesinclude letting it naturally stick to a surface (termed adsorption),making a charge-charge bond with the surface (e.g., an ionic bond), andattaching it via an even stronger, more permanent chemical linkage suchas a covalent bond. Heparin has been applied as a thin coating ofpolymers adsorbed to a surface by dipping the surface into a solution ofheparin or drying the heparin onto the surface. Heparin has a negativecharge and has been exposed to surfaces that have a positive charge sothat it remains there via a charge-charge interaction. Photoactivatedchemical groups have been put onto heparin so that the heparin is putclose to the surface, the surface is bathed in light, and thephotoactive groups make permanent covalent chemical bonds between theheparin polymer and the surface. Similarly, heparin has been chemicallyattached to monomers that have then been reacted with the surface.

Patent families and patent applications that describe the use of heparininclude U.S. Pat. No. 6,127,348, which include descriptions ofcross-linked alginate and certain other polysaccharides as compositionsuseful for inhibiting fibrosis. U.S. Pat. No. 6,121,027 includesdescriptions of decorating heparin with a photoactive cross linkingchemical group. Application PCT GB9701173 and U.S. Pat. No. 6,096,798include descriptions of heparins with monomers used to make polymers.U.S. Pat. Nos. 5,763,504 and 5,462,976 include descriptions ofglycosaminoglycans derivatized with photoactive groups and cross-linkedthereby. U.S. Pat. No. 6,060,582 includes descriptions of macromere witha water soluble region, a biodegradable region, and at least twofree-radical polymerizable regions. Other patents include descriptionsof a polysaccharide reacted with other polymers, decorated with apolymerizable group, and/or reacted to form a coating on a surface;including U.S. Pat. Nos. 5,993,890; 5,945,457; 5,877,263; 5,855,618;5,846,530; 5,837,747; 5,783,570; 5,776,184; 5,763,504; 5,741,881;5,741,551; 5,728,751; 5,583,213; 5,512,329; 5,462,976; 5,344,455;5,183,872; 4,987,181; 4,331,697; 4,239,664; 4,082,727; and Europeanpatents 049,828 A1 & B1.

Despite many years of research in the areas of polysaccharides,hydrogels, and blood-contacting materials, the need for betterimplantable synthetic materials that cause little or no unfavorablereaction from a patient's body remains acute. In particular, there is agreat need for a medically useful small diameter vascular graft made ofsynthetic materials.

SUMMARY OF THE INVENTION

The present invention meets all of these needs by providing embodimentsthat include synthetic materials that successfully combine theadvantages of polysaccharides and hydrogels in a medically useful mannerso that they may be used for devices, including small diameter vasculargrafts. Other embodiments of the invention include hydrogels made ofpolysaccharides and materials and methods for making such hydrogels, aswell as products that incorporate such hydrogels. One embodiment of theinvention is a hydrogel made of a polysaccharide, for example heparin.Another embodiment is a hydrogel made by polymerizing heparin macromers.Another embodiment is a small diameter vascular graft made bypolymerizing heparin macromers around a tube, for example a polyestertube.

An embodiment of the invention is a medical apparatus havingpolysaccharide macromers polymerized into a three-dimensionalcrosslinked hydrogel that makes a hollow cylinder; with the cylinderbeing formed during polymerization of the polysaccharide macromers. Ahollow cylinder is essentially equivalent to a tubular structure and maybe made out of any material, e.g., metals, plastics, ceramics, having avariety of properties, including e.g., rigid, compliant, and elastic.

Another embodiment of the invention is a biocompatible encapsulation foran inert medical device. The encapsulation has polysaccharide macromerspolymerized into a three-dimensional crosslinked hydrogel thatencapsulates the medical device.

Another embodiment of the invention is a medical apparatus made of amaterial that has a heparin macromers polymerized into athree-dimensional crosslinked hydrogel that forms a hollow cylinder thatis not covalently bonded to another material.

Another embodiment of the invention is a pollyvinylpyrrolidone macromer.Another embodiment is a biocompatible coating system that haspolyvinylpyrrolidone macromers polymerized into a three-dimensionalcrosslinked material that contacts a medical device and thereby formsthe coating.

Another embodiment of the invention is a polysaccharide polymer of atleast two polysaccharide macromers polymerized together. The polymer ispreferably in an isolatable form. The macromers may have polymerizablemoieties such as polyhydroxyethylmethylacrylates, methyl methacrylates,methacrylates, acrylates, photopolymerizable monomers, monomers withhydroxyl groups, monomers with glycerol groups, monomers withpolyoxyalkylene ether groups, monomers with polypropylene oxide groups,monomers with vinyl groups, monomers with zwitterionic groups, monomerswith silicone groups, monomers having sulphate groups, monomers havingsulphonate groups, and heparin monomer.

Another embodiment of the invention is a method of making a medicalapparatus from a material that includes a plurality of polysaccharidemacromers by polymerizing the macromers into a three-dimensionalcrosslinked hydrogel that defines a hollow cylinder, wherein thecylinder is formed during polymerization of the polysaccharidemacromers.

Another embodiment of the invention is a method of encapsulating aninert medical device by polymerizing a plurality of polysaccharidemacromers into a three-dimensional crosslinked hydrogel thatencapsulates the medical device.

Another embodiment of the invention is a method of making a medicalapparatus by polymerizing heparin macromers into a three-dimensionalcrosslinked hydrogel and thereby making a hollow cylinder having anexterior, wherein the cylinder is formed during polymerization of theheparin macromers and the exterior is not covalently bonded to anothermaterial.

Another embodiment of the invention is a method of makingpolyvinylpyrrolidone polymers from polyvinylpyrrolidone macromers.Another embodiment is coating a medical device with polyvinylpyrrolidonemacromers by polymerizing the macromers into a three-dimensionalcrosslinked polyvinylpyrrolidone material, and applying a coating thatincludes the crosslinked polyvinylpyrrolidone material onto the medicaldevice.

Another embodiment of the invention is a method of making apolysaccharide polymer by obtaining or making polymerizablepolysaccharide macromers, synthetically polymerizing the macromers witheach other to form a group of polymers having an average length of atleast two macromers per polymer, and isolating the polymers.

Another embodiment of the invention is a method of making a coating on amedical device by providing a group of polysaccharide polymers having anaverage length of at least two macromers per polymer, putting thepolymers in a solvent to make a mixture, and contacting the medicaldevice with the mixture.

Embodiments of the invention include a material made of a hydrogel thatpreferably has at least 5% polymerized polysaccharide macromers by dryweight. The hydrogel is preferably covalently cross-linked such that thehydrogel remains intact in water and preferably contains at least 30%water by total weight when hydrated. The polysaccharide macromers arepolymerizable while in a solution or in a suspension. Normalpolymerization techniques, including free-radical, addition, andcondensation polymerization, may be used to polymerize thepolysaccharide macromers.

One product incorporating the present invention is a tubular member withits inner wall and outer wall covered with a hydrogel of the inventionas described herein, i.e., the tubular member is “encapsulated” by thehydrogel. A preferred macromer formulation is made from heparin, theterm heparin including all molecular weights of heparin, heparansulfate, heparan sulfate proteoglycans, fragments thereof, and/orderivatives thereof. The preferred embodiment of the heparin hydrogel isat least 80% heparin by dry weight. The tubular member preferably has adiameter of less than approximately 6.0 mm when the hydrogel is hydratedand blood is flowing though the tubular member. The tubular member maybe a simple plastic extrusion or a stent, but the preferred embodimentis a knitted or woven fabric substrate. The fabric substrate ispreferably pre-coated to enhance the integrity and adhesion of theencapsulant and/or improve the non-thrombogenic or anti-thrombogenicproperties or both of the encapsulant.

In an embodiment of the tubular member, the tubular member is pre-coatedwith a very thin layer of the hydrogel containing non-thrombogenic oranti-thrombogenic properties or both, the layer being applied to coverthe components of the tubular member. In the case of the knitted orwoven tube, these are the individual strands from which the fabric ismade. The tubular member preferably has a low porosity such that bloodleakage is not a paramount concern.

In another embodiment, a porous tubular member, such as one made from afabric, is used as the tubular member. The fabric tubular member ispre-coated with a polymeric material in order to prevent blood leakage.The coated fabric tubular member is then further coated with thehydrogel containing non-thrombogenic or anti-thrombogenic properties orboth. This structure imparts an extremely thin, flexible, and compliantwall that can serve as a vascular prosthesis, especially in the contextof a small diameter vascular graft.

Another embodiment is a tissue engineering matrix made from apolysaccharide hydrogel. A tissue engineering matrix is, for example, athree-dimensional material that serves as a scaffold for cellularinvasion or a nerve growth matrix. Examples of tissue engineeringmatrices include matrices for making cartilaginous body parts such asears or joint cartilage; ligaments; scaffolds for breast tissueinvasion; liver matrices; and tissue engineered blood vessels.

The present inventors have also recognized that there is a need to usebetter organic solvents to dissolve polysaccharides, including heparin.The use of better organic solvents allows scientists to use chemistriesand chemical techniques that are more powerful than those that areconventionally used. These techniques improve the cost, quality, andefficiency of conventional techniques for making materials frompolysaccharides and enable better materials to be made.

Embodiments of the invention include the use of low dielectric organicsolvents and/or low boiling-point solvents for polysaccharidechemistries to make derivatives of polysaccharides, including attachingmonomers to polysaccharides to make polysaccharide macromers and the useof these improved solvents for making polysaccharide hydrogels frompolysaccharide macromers or polysaccharide polymers. Further, theinvention may include steps for using salts to decomplex the quaternaryammonium-heparin complex.

An embodiment of the invention is a method of making a polysaccharidemacromer, for example from heparin. The polysaccharide is reacted with aquaternary ammonium salt to form a polysaccharide-quaternary ammoniumsalt complex and then dissolved in an organic solvent with a dielectricconstant less than the dielectric constant of DMSO and/or in an organicsolvent with a boiling point less than DMSO. Thepolysaccharide-quaternary ammonium salt complex may be reacted with achemical such as a monomer to form useful derivatives. Thepolysaccharide-quaternary ammonium salt complex may then be treated withanother salt to remove the quaternary ammonium salt.

The invention optionally includes steps of using a vacuum to remove theorganic solvent of the invention from the polysaccharide, derivatizedpolysaccharide, or complexes of the polysaccharide. The vacuum removalis preferably performed at room temperature without adding heat.Alternatively, heat may be applied to evaporate the solvent, preferablyenough heat to raise the temperature of the solvent to its boiling pointwithout denaturing the heparin such that its biological activity issubstantially reduced, a temperature that may vary according to thesolvent used but typically being a temperature of less thanapproximately 100 degrees Centigrade and preferably less than 70 degreeCentigrade. Alternatively, a mix of vacuum and heat may be used.

A preferred embodiment of the invention uses an organic solvent that hasa boiling point at atmospheric pressure and a dielectric constant thatare less than conventionally used organic solvents. A more preferredembodiment has a boiling point of less than approximately 115 degreesCentigrade and a dielectric constant that is less than that of DMSO. Amore preferred embodiment uses an organic solvent that has a boilingpoint of less than approximately 70° C. at atmospheric pressure and adielectric constant that is less than that of DMSO.

The invention includes polymerizing the polysaccharide macromer in anorganic solvent of the invention to make a polymer of at least twomacromers. The macromers may be the same or different, to thereby make ahomopolymer or a copolymer. The invention includes making a hydrogelfrom the polysaccharide macromer and/or polysaccharide macromer,preferably in an organic solvent of the invention.

BRIEF DESCRIPTION OF THE FIGURES

FIG. 1. is a longitudinal sectional view of an embodiment of theinvention having a fabric graft encapsulated within a hydrogel;

FIG. 2. is a side elevational view of the structure depicted in FIG. 1;

FIG. 3. depicts a longitudinal sectional view of an alternativeembodiment of the invention having a fabric graft coated with a hydrogelof the invention;

FIG. 4. depicts a longitudinal sectional view of another alternativeembodiment of the invention having a double-coated fabric graft; and

FIG. 5 depicts a longitudinal sectional view of another alternativeembodiment of the invention like the embodiment of FIG. 4 except thatthe fabric graft tube has been inverted.

FIG. 6 depicts a partial view of an alternate embodiment of theinvention having a plastic surface coated with a hydrogel.

FIG. 7 depicts a cross-sectional view of a hydrogel of the inventionbeing formed on a mandrel.

FIG. 8 is a sectional view of FIG. 7.

FIG. 9 depicts a reaction scheme for making and using polysaccharidemacromers.

FIG. 10 depicts an alternative reaction scheme for making and usingpolysaccharide macromers.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

Synthetic materials implanted into the soft tissue of a patient elicit arange of undesirable reactions that are typically categorized as acuteinflammation, chronic inflammation, the formation of granulation tissue,foreign body reaction, and fibrosis (Ratner et al., 165-173,Biomaterials Science, 1996 Academic Press). Other reactions arepossible, such as an immune system response or systemic toxicity andhypersensitivity. At a materials-blood interface, the blood coagulationmechanism can be activated to cause local clotting and downstream eventssuch as complement activation or generation of blood clots (Ratner etal., 193-199). These reactions impact how synthetic articles, includingvascular grafts, are designed and manufactured.

Current medical practices for the use of synthetic articles as vasculargrafts are essentially restricted to the use of polyester andpolytetrafluoroethylene tubes as replacements for large-diameter bloodvessels to treat patients with certain types of vascular disease.

Vascular disease takes many forms, but one of the most common isstenosis, where an artery becomes narrowed by build up of plaque.Atheroscherotic Stenosis is a condition where the artery becomeshardened and less flexible by the process of calcification. Thecondition is often accompanied by a build up of tissue or plaque on theinside wall of the lumen of the artery. This build up causes the lumenof the artery to narrow and restrict the passage of blood.

Total occlusion of an artery may occur when a clot of blood (a thrombus)lodges inside an artery, at a narrowing. Such occlusions can prove fatalif the artery is in a critical position; for example, coronarythrombosis causing a heart attack, or cerebral thrombosis causing astroke.

Even if there is not total occlusion, a restriction of the flow of bloodcan cause severe problems in limbs and organs downstream from thestenosis, due to starvation of the oxygen and nutrients supplied by theblood. A very common example is the reduction of flow in the lower leg,which may lead to claudication and eventually to gangrene and loss of alimb.

A higher profile example of stenosis is observed when the diseaseaffects the coronary arteries, which supply blood to the muscle of theheart. Stenosis, or occlusion due to thrombosis, can lead to aninfarction due to parts of the muscle tissue of the heart dying(necrosis). That is, a myocardial infarction or heart attack.

Another common form of vascular pathology is an aneurysm. An aneurysm isa condition where the wall of the artery weakens and dilates to form aballoon-like swelling. An aneurysm is a weakening of the artery wallleading to dilation. This dilation can develop so that the arterial wallis too thin and weak to withstand the pressure of the blood. A burstaneurysm causes severe hemorrhage, and can be fatal. One example is theso-called AAA (triple A) or abdominal aortic aneurysm; rupture of anartery having this defect is almost always fatal.

Conventional vascular surgery techniques have made the surgicalreplacement of diseased arteries commonplace. The use of autologousgrafts where a non-essential vein from a person's own blood vessels isused as a replacement artery is the oldest form of vascular grafting,and is still used today especially for coronary bypass procedures. Apatient, however, may not have enough autologous donor vessels to fillthe needs of the replacement surgery. Further, it is desirable thatthere be an alternative to the loss of a functioning blood vessel. Itwas the development in the 1960s of synthetic fabric prostheses, whichled to the range of products available to the vascular surgeon today.

Conventional synthetic vascular grafts are woven or fabric seamlessfabric tubes, which are used as a direct replacement for a section ofdiseased artery. Various materials have been tried, but the mostsuccessful is polyester, which is the only material now used byclinicians for fabric grafts. Polyester is a very bio-stable materialand, although slightly thrombogenic, is reasonably well tolerated in usefor larger diameter arteries (approximately 6-mm diameter and above).

When implanted, a vascular graft causes the body to react and generate ablood clot layer around its inner perimeter. This blood clot may beaccompanied by tissue growth. If such tissue growth is not securelyanchored in place, there is a danger of a loose blood/tissue clot orembolus being formed. The flow of blood may carry the embolus downstreamuntil it reaches a narrower part of the artery where it may cause ablockage or occlusion.

In order to ensure that tissue growth remains anchored to the graft,fabric grafts are made slightly porous so that the tissue grows into thepores of the fabric and is firmly attached. This has an added advantagethat the tissue encapsulates the fabric graft and covers it, using thegraft as scaffolding for new growth. In general, grafts with higherporosities, and especially flexible and elastic fabric grafts, healbetter and generally perform better than low porosity woven grafts. Itis because of this effect that the so-called “velour” grafts have beendeveloped. Velour grafts have textile filaments raised up on the surfaceof the fabric.

The need for porosity creates an initial problem for the surgeon since ahigh porosity graft will leak if simply implanted with no specialpreparation. After implantation, tissue growth fills in the pores in thefabric and therefore renders the graft “blood tight”. The problem is tomake the graft blood tight during the initial surgery and for theimmediate few hours afterwards. The original approach taken was to “preclot” the graft before implantation. A small amount of blood was takenfrom the patient before the operation, and the graft was soaked in thisblood so as to fill all the pores of the fabric with clots. The porousstructure ensured that the clotted blood was firmly attached. Thisprocedure was, however, time consuming, and it was and is very difficultto properly pre-clot those materials with high porosities. Because ofthese problems, the soft and very compliant fabric grafts, which healbetter than stiff woven grafts, were not considered suitable for areasof high pressure such as thoracic arteries near the heart.

Subsequent to the development of fabric grafts, a new development wasintroduced. The invention of the graft made from polytetrafluoroethylene(PTFE). PTFE is a material that is generally well tolerated in the body.A chief advantage with a PTFE graft is that it does not require preclotting. However, PTFE material does not heal as well as warp fabricgrafts. Modern PTFE prostheses are made of expanded material in order toimitate the cellular structure of fabric grafts, but this is onlypartially satisfactory.

Fabric grafts are now available which do not require pre-clotting. Theyare coated with a bio-absorbable material such as gelatin or collagen.The coating material is gradually adsorbed and tissue grows to replaceit. Such grafts are not only simpler to use because there is norequirement for pre-clotting, but they also tend to give better healingand performance.

All of these conventionally used grafts have a common disadvantage inthat they cannot be used for small diameter applications. Syntheticvascular grafts with a diameter of less than approximately 6 mm will notremain functional over a clinically significant period of time.

The reason for the failure of conventional small diameter vasculargrafts is thought to be related to the flow rate and type of blood flowwithin the graft. With a large inner diameter graft, there is a highvolume of blood passing though the vessel. Any small tissue growth onthe inside wall is insignificant and does not disrupt flow. With smallinner diameter grafts, a small tissue build up is more significant inrelation to the overall diameter of the vessel. A small build up willcause flow turbulence, which, because of the already low volume flow,tends to cause even more tissue growth, which leads to stenosis.

Conventional medical clinicians urgently need a small diameter syntheticgraft to cope with a range of small caliber replacement requirements.The only procedure conventionally available is an autologous transplantof the saphenous vein in the leg. The problem with this approach is thatthe amount of graft material available is very limited. In addition, theremoval of the saphenous vein causes severe discomfort to the patient.One of the most important of these small diameter applications is thereplacement or bypass of coronary arteries in cases of coronarystenosis. Another use is for the replacement of stenosed or occludedinfragenticulate arteries.

In order to replace the smaller caliber vessels using a small diametervascular graft, the implant should be highly biocompatible and notthrombogenic. The biocompatibility should be such that little tissuegrowth is stimulated and the blood and body accept the prosthesis in asimilar way that they accept the natural artery.

Many polymers have been tested for the small diameter vascular graftapplication. A polymer is a single molecule built up by the repeatedreaction of molecules that may be referred to as monomers. Monomers aremolecules that may be reacted with other monomers to make a polymer. Forexample, a methylmethacrylate monomer represented by “A” may be used tomake polymethylmethacrylate, which would be the polymer “AAAAA”. Apolymer made of monomers A and B could have a random structure ofABAAABAABA and could be called a polymer or a copolymer of A and B.Polymer AAAAA and BBBBB could be joined to form copolymer AAAAABBBBB. Amolecule that is derivatized is a molecules that has been chemicallychanged; a molecules that has been decorated is a molecules that has hada chemical unit attached.

A macromer, as used herein, is a monomer or a polymer that ispolymerizable and is a convenient term for referring to monomers orpolymers that have been decorated with a monomer or used as decorations.Polysaccharide macromers include multiple polysaccharides that arepolymerizable. For example, several macromers may be polymerizedtogether to form a larger polymerizable group that is a macromer. Orseveral polymers, e.g., polysaccharides, may be joined together anddecorated with polymerizable groups to form a macromer. Thus a heparinmacromer is a heparin molecule that is polymerizable, for example afterbeing decorated with a monomer. The term polymer, as used herein,includes oligomers and chains of at least two monomers in length.Polymers can be assembled into larger materials; for example, manypolymers may be linked together to form a hydrogel.

A variety of hydrogels are blood compatible. Examples that arecommercially useful are made from polyhydroxyethylmethylacrylate(PHEMA), polyacrylamides, polyacrylic acid, N-vinyl-2-pyrrolidone (NVP),methacrylic acid, methyl methacrylate, and maleic anhydrides, each ofwhich have been proven to be polymerizable from monomers. Furtherexamples of biocompatible materials include hydrogels made frompolyvinyl alcohol, methacrylates in general, acrylates in general,polyethylene glycol (PEG), hyaluronic acid, and alginates.

The term hydrogel, as used herein, is a cross-linked material that canabsorb or imbibe a water and is produced by the cross linking of one ormore monomers or polymers. The cross-links in a hydrogel may be theresult of covalent bonding or of association bonds, for example hydrogenbonds, charge-charge interactions, or strong van der Waals interactionsbetween chains (Ratner et al., pages 60-64). A hydrogel cannot besuspended in water nor does it dissolve in water; instead, it remainsintact water. For example, a material can shrink or swell and stillremain intact without dissolving. For example, a contact lens shrinksand swells in water but does not dissolve therein.

Hydrogels have been used for a myriad of applications, such asartificial tendon materials, wound-healing bioadhesives, wounddressings, artificial kidney membranes, articular cartilage, kneecartilage replacement, artificial skin, maxillofacial and sexual organreconstruction, tissue engineering scaffolds, and vocal cord replacementmaterials. There are many types of hydrogels known to those skilled inthese arts, such as aerogels, xerogels, equilibrium-swollen hydrogels,solvent-activated hydrogels, and swelling-controlled relapse hydrogels(e.g., Ratner et al., page 60-64). These types of hydrogels may be usedwith certain embodiments of the invention. Moreover, all of the uses forhydrogels described in this application may be used in certainembodiments of the invention.

Polysaccharides are polymers made from monomers that are sugars.Alginate is a polysaccharide. Glycosaminoglycans are a subcategory ofpolysaccharides that are made from repeats of disaccharide units.Glycosaminoglycans include hyaluronic acid, chondroitin sulfate,dermatan sulfate, heparan sulfate, heparin, chitin, chitosan, andkeratan sulfate. Polysaccharides and glycosaminoglycans can also befound in a related proteoglycan form; proteoglycans are polysaccharidesunified with a protein. Methods, compositions, and uses as describedherein for polysaccharides are applicable to proteoglycans,glycosaminoglycans, and natural or synthetic derivatives or fragmentsthereof.

Heparin is a polysaccharide polymer with an important property: itinterferes with key molecules in the blood clotting mechanism such thatthe blood will not clot. Coating the inside of a tube with heparin tendsto increase the amount of time that the tube remains open to blood flowbut, to date, small diameter vascular grafts coated with heparin havefailed to resist blockage by blood coats for a clinically useful lengthof time. Many researchers have tried to use heparin to coat the insideof vascular grafts. Heparin is a molecule that belongs to a group ofmolecules called polysaccharides that are molecules made fromcombinations of smaller molecules called sugars. Polysaccharides arenaturally-occurring polymers but the present invention also contemplatessynthetic, derivitized, man-made, and semi-synthetic polysaccharides.

Heparin has been applied to materials in many ways. General strategieshave included adsorption, making charge-charge bonds with the surface,covalent immobilization, and release from a surface. Photoactivatablechemical groups have been put onto heparin so that if heparin is putclose to the surface and the surface is bathed in light to make thephotoactive groups make permanent covalent chemical bonds with thesurface. Similarly, heparin has been reacted with polymerizable monomersthat have then been reacted to achieve covalent bonding to the surface.

Despite the great amount of research effort that has been expended inthe field of biomaterials, including research with hydrogels, heparin,and polysaccharides, there is a continued need for improvedbiomaterials. Biomaterials that may be used in blood-contactingapplications are especially required.

The invention provides embodiments that include an improved biomaterialthat successfully combines the advantages of polysaccharides andhydrogels. One embodiment of the invention is a hydrogel made ofpolysaccharides. Another embodiment is a linear or multi-armedpolysaccharide or polyvinylpyrrolidone polymer that is absorbable to asurface. Multi-armed means a soluble polymer that is branched orcross-linked. The hydrogels and the coatings may be used for the manyapplications for which a hydrogel may be used, e.g., as alreadydescribed. There are other uses for hydrogels and the linear ormulti-armed polymers that include coatings for stents, cathetercoatings, cardiac valves or leaflets, cartilage replacement, replacementknee cartilage, organ scaffolds, lumbar disks, cell encapsulation, woundhealing, nerve guides or tubes, and postoperative adhesions. A hydrogelmay be used to make, coat, or encapsulate such devices. A coating may beused to such devices to improve their performance.

Another hydrogel use is for a wound dressing for large, shallow woundson animals so that a scab does not form over the wound but preventsblood clotting at the surface, thereby preventing scar formation.

The hydrogel may be made by mixing polysaccharide macromers, e.g.,heparin macromer, with other macromers or monomers to make a mixture.The mixture is poured into a mold and polymerized. After polymerization,the mold is removed and the polymerized macromers/monomers are hydrated.Various shapes, e.g., sheets, tubes, spheres, rods, may be formed byusing suitable molds.

The resultant shape is not covalently or otherwise bonded to othermaterials: the exterior and luminal surfaces are “free”. The freesurfaces are not attached to other surfaces. A free surface may bedecorated with moieties, e.g., drugs, polymers, and other agents. Suchdecorations do not cause the free surface to thereby be attached toother materials. Further, the resultant shape is formed during thepolymerization process. This shape-forming process is distinct fromprocesses that build up a coating on the inside of, e.g., a tube, inpart, because the coating on the tube is essentially inseparable fromthe tube, especially if it is ionically or covalently bonded thereto.

Moreover, the prior art methods of applying a coating to a tube andbuilding up the coating is not equivalent to the present process. First,the prior art coating procedure does not create the tubular shape. Butpolymerizing macromers of the invention into a tubular shape during thepolymerization process does create the tubular shape. Making a shapeduring a polymerization process is difficult because the polymerizationreaction must be effective enough to make a solid material, with theeffectiveness depending on polymerization variables known to thoseskilled in these arts. For example, an effective process requires usinga macromer that can be provided in sufficient concentration. Not allmacromers are sufficiently soluble to be present in solution with a highenough concentration to make a solid. The polymerization mixture musthave be crosslinkable for crosslinks to form. The kinetics of thepolymerizable groups must be suitable.

Many prior art processes have not overcome these limitations and insteaddry polymerizable groups onto a surface and then crosslinking them. Thedrying step results in drastically different structures than those madewith polymerization from a solution (or a melt). If proteins orpolysaccharides are dried, they aggregate and form clumps on themolecular level. In contrast, polymerization from a solution givesstructures that are not aggregated but instead have a network ofunaggregated polysaccharides. Since the macromers are not aggregated,the density of materials polymerized from solution can be lower.Moreover, a true polymerization process may take place wherein thepolymerizable groups react with each other to form a polymeric backbone.In contrast, dried solutions have little mobility and the polymerizablegroups react with the chemical structures closest to them instead ofreacting with other polymerizable groups.

In short, chemical crosslinking is not equivalent to polymerization. Thepresent materials and methods provide for polymerization methods and forpolymerized materials as opposed to coatings built up on surfaces,gelled structures, and merely aggregated, chemically crosslinked, orsurface-immobilized materials. The advantages of polymerization arenumerous and well known to ordinary artisans.

One advantage of a polysaccharide covalently polymerized with a hydrogelis that the polysaccharide may be stably incorporated into the hydrogelso that it is not released over time. This stability is useful for along-term implant because the hydrogel would otherwise dissipate overtime and eventually fail. The heparin hydrogels of the present inventionare hypothesized to function by reversibly binding antithrombin III. Theantithrombin III is bound by the heparin and thereby changes its shapeso that it reacts with and inactivates both thrombin and Factor X(a),which are key enzymes required for blood to clot. The antithrombin IIIis hypothesized to stay on the heparin hydrogel temporarily so that itattaches, reacts with thrombin and Factor X(a), and departs back intothe bloodstream so that a new antithrombin III molecule may be bound tothe heparin.

Another advantage of the polysaccharide hydrogel is that it can be madeas a thick film. Thick films may be handled by surgeons, grasped withtools such as forceps, punctured by sutures, and suffer scratches anddamages to their surface without losing their favorable blood-contactingproperties. The thickness of the film and the three-dimensionalstructure of the film allows it to suffer minor damage while continuingto cover the surface with heparin molecules. In contrast, damage to athin coating or a synthetic material that has been merely reacted with apolysaccharide can entirely remove the polysaccharide and expose theunderlying material to the body. For example, a surgeon that usesforceps to firmly grasp a plastic tube covered with a layer of heparinthat has been reacted with the tube's surface might accidentally scratchthe tube and remove the heparin thereby exposing the underlying plasticmaterial of the tube. In contrast, a surgeon might accidentally scratcha plastic tube encapsulated with a thick film heparin coating of thepresent invention but would not thereby expose the underlying plasticbecause a scratch in the thick film would expose only more of theheparin hydrogel.

Further, the thickness of the hydrogel film is hypothesized to minimizeblood contact with the synthetic material that the thick film isencapsulating. Blood or its components must penetrate through the entirethickness of the hydrogel prior to reacting with the encapsulatedsynthetic material. In contrast, a thin coating, especially a coating ofa few molecules' thickness, presents a shorter distance between theblood and the encapsulated synthetic material. This distance isimportant because the efficiency of surface chemistry reactions used forconventional techniques is hypothesized to typically provide a surfacecoverage of less than 100%, i.e., not every space on a surface coatedwith heparin is completely covered with a heparin molecule. In contrast,a thick film of the present invention provides more than 100% coveragebecause any molecule that would react with the surface must pass througha thick coating that has a thickness of many molecules.

One embodiment of the invention is a tubular member encapsulated by apolysaccharide hydrogel. The encapsulation may be achieved by a numberof processes, one such process being placing tubular member into a moldand charging the mold with the desired formulation of polysaccharidemacromers and polymerized by conventional techniques at room or atelevated temperature and/or by electromagnetic radiation. A materialthat encapsulates a member is termed an encapsulant.

Certain embodiments of the invention involve polymerization processes,including polymerization of polysaccharide macromers. The polysaccharidemacromers may be polymerized using conventional techniques, for exampleusing initiators, carriers, accelerators, retardants, viscosifiers,and/or cross-linkers. The polysaccharide macromers may also becopolymerized with and other monomers and/or macromers and/orpolymerizable polymers. In another embodiment, a tubular member iscovered with a mixture of polysaccharide macromers and/or monomersand/or polymers. The polymerizable groups are subsequently polymerizedto form a crosslinked hydrogel. It may be desired to placed the tubularmember on a rotating mandrel and apply the mixture so as to ensure thatthe tubular member is encapsulated uniformly.

The polysaccharide macromers and polymers made thereof may also becopolymerized, blended, mixed, and/or cross-linked with other monomersand/or macromers, and/or polymers including engineering polymers, bloodcompatible polymers, hydrogel polymers, natural polymers (e.g.,deoxyribonucleic acid, polysaccharides, and proteins and bioactivefragments thereof) and/or fillers. The type of initiation is not limitedand may include thermal, X-ray, ultraviolet, infrared, visible light,free radical, addition, sonic, and condensation initiation.

Monomers for mixing with the polysaccharide macromers can include, butare not limited to, monomers with hydroxyl groups (e.g., hydroxyethylmethacrylate), monomers with glycerol groups (e.g., glycerolmonomethacrylate, glycerol dimethacrylate, glycerol trimethacrylate),monomers with polyoxyalkylene ether groups (e.g., polyethylene glycolmethacrylate, polypropylene glycol methacrylate), monomers with vinylgroups (e.g., N-vinyl pyrrolidone), monomers with zwitterionic groups(e.g., 2-methacryloyloxyethyl-2-(trimethyl ammonium)phosphate, monomerswith silicone groups (e.g., methacryloxypropyltris(trismethyl-siloxy)silane and other silicone methacrylate oracrylates), monomers having sulphate groups (e.g., vinyl sulphonicacid), monomers having sulphonate groups (e.g., ammoniun sulphatoethylmethacrylate), heparin monomer as cited in the patent PCT GB9701173 andU.S. Pat. No. 6,096,798, which are hereby incorporated herein byreference.

Polymers for mixing, blending, and/or copolymerization withpolysaccharide macromers include derivatized polymers, for example,derivatized polyoxyalkylene ether groups (e.g., polyethylene oxideterminating in hydroxyl group, carboxylic groups and/or isocyanategroups and polypropylene oxide terminating in hydroxyl group, aminogroups, carboxylic groups and/or isocyanate groups), polyvinylpyrrolidone functionalized with methacrylate groups, methacrylateterminating dimethysiloxone, vinyl terminating dimethylsiloxone,polyurethane terminating in isocyanate, polyester terminating inisocyanate and also other polymers that can be derivatized withmethacrylate, acrylate, isocyanate, carboxylic acid, amino, hydroxyland/or vinyl groups.

Polymers mixed, blended and/or copolymerized with polysaccharidemacromers may also be used to enhance the viscosity of the mixtureformulation for the application to the rotary mandrel and hence thepolymerization of the mixture formulation. This technique may be used togive the encapsulated tubular member a homogenous and smooth surface, afeature that enhances vascular prosthesis biocompatibility. Theencapsulated tubular member and hence the biocompatible vascularprosthesis has a smooth surface containing non-thrombogenic oranti-thrombogenic properties or both and preferably has a water contentranging from about 30% to about 90%.

In a first preferred embodiment of the invention, a tubular member isencapsulated by the hydrogel, e.g., as in FIGS. 1 and 2, the hydrogelhaving cross-linked polymers made from polysaccharide macromers andcopolymerized with monomers from at least three of these classes, asdiscussed in patent application PCT GB97 01173, U.S. Pat. No. 6,096,798:(a) monomers having sulphate groups, (b) monomers having sulphonategroups, (c) monomers having sulphamate groups, (d) monomers havingpolyoxyalkylene ether groups and (e) monomers having zwitterionicgroups. The polysaccharide macromers are preferably heparin macromers.Hydrogel encapsulation of the tubular member is performed by placing thetubular member into a mold, adding a macromer and/or monomer formulationand then polymerizing to make a hydrogel. The monomer constituents mayvary from 10% to 90% by weight and are preferably polymerized with abifunctional monomer, e.g., ethylene glycol dimethacrylate. Thisformulation provides the prosthesis with a smooth surface, prevents theleakage of blood, is non-thrombogenic and/or anti-thrombogenic, and hasa water content ranging from 30% to 90% when hydrated. When hydrated,the biocompatible vascular prosthesis is soft and pliable so it will notcompromise the mechanical properties of the prosthesis.

FIG. 1 depicts a longitudinal cross section of a small diameter vasculargraft 10 according to a first preferred embodiment of this invention.FIG. 2 shows an end view of a radial cross section of FIG. 1. The loopsof synthetic material 12 are completely encapsulated within hydrogel 14.This embodiment is suitable for porous members that allow passage of themacromer and/or monomer constituents through the pores of the tubularmember, e.g., a fabric, in order to provide binding between the innerand outer faces of the encapsulant.

An alternative to the first preferred embodiment of the invention usesapproximately the same materials and methods but a tubular blood vesselmember 13 is used which has sufficiently low porosity so that bloodleakage is not a consideration. e.g., a tightly knitted or woven fabric,or a plastic extrusion. With this kind of tubular member, the basicporosity is low. Referring to FIG. 3, instead of encapsulating tubularmember 13, it is covered with a hydrogel layer 14 of polysaccharidemacromer and/or monomers. In the case of a fabric tubular member, thehydrogel coats the individual yarns and fibers. At the same time, thehydrogel may be used to cover the pores completely or partially. Theloops of knitted fabric 13 are coated by hydrogel 14. This embodiment issimpler to make than the encapsulation process as the covering isapplied by dip process, spraying technique or by other conventionalprocess.

A second alternative to the first preferred embodiment of the inventionuses approximately the same materials and methods but uses a poroustubular member 15, such as one made from a fabric, which is pre-coated(a primary coating layer 16) with a polymeric material in order toprevent blood leakage, as the layer provides a strong “leak proof”security layer (FIG. 4). The leak proof layer is essentially impermeableto blood, meaning that it generally prevents flow of blood in surgicalapplications. Alternatively, the barrier may be permeable to blood, andcontrol the flow of blood. Primary coating layer 16 may be made from anumber of flexible polymers, e.g., silicone polymer, polypropylene,polyester, polyurethane, polytetrafluoroethylene (PTFE) or elastomericpolymer such as silicone rubber. The total composite is then furthercoated with hydrogel 14 containing polysaccharides. This structureimparts an extremely thin, flexible, and compliant wall to vascularprosthesis 10. Each of the coating processes may be applied by dipprocess, spraying technique or by other conventional coating process.FIG. 4 depicts a longitudinal cross section of a knitted tubular membercoated according to this embodiment. The fabric loop of porous tubularmember 15 is inside primary coating 16, which in turn is inside hydrogel14.

Referring to FIG. 5, another embodiment of the invention takes advantageof the manufacturing processes of woven vascular prostheses. A tubularmember made from knitted fabric has a different surface texture on theinternal face and the external face. The technical face of the fabrictends to be smoother than the technical back. In the case of the knittedtubular member, the technical back is on the internal face of thefabric, whereas the technical face is on the external face. When used asa vascular prosthesis, the knitted tubular member can be inverted sothat the smoother face is on the internal face of the prosthesis. Inaddition, the internal face has very shallow grooves, which runlongitudinally. These grooves may help to smooth the flow of the bloodand reduce turbulence.

FIG. 5 shows a schematic longitudinal cross section of a porous tubularmember 15 made of a knitted material that is coated with the fabricinverted. The fabric loop of the porous tubular member 15 is inside theprimary coating 16, which in turn is inside the hydrogel 14. A tubularvascular vessel made from knitted fabric has a different surface textureon the internal face and the external face. The technical face of thefabric tends to be smoother than the technical back. In the case of theknitted tubular member, the technical back of the fabric is on theinternal face of the vessel, whereas the technical face is on theexternal face of the vessel. The fabric loop of the porous tubularmember 15 is inside the primary coating 16, which in turn is inside thehydrogel 14.

A most preferred tubular member is a warp knitted or weft knittedseamless fabric tube. In the warp knitted form, the most preferredstructure is reverse locknit, but tricot can also be used. The fabrictube can also be weft knitted or woven. The tube may be continuous or itmay be bifurcated in order to fulfill the needs of a graft designed toreplace the Aorto-Iliac bifurcation.

FIG. 6 shows a portion of a vascular graft 10 with surface 18 with athick film of hydrogel 24. The thickness of film 24 protects surface 18from being exposed as a result of damage caused by handling the vasculargraft 10. The film is preferably at least 25 μm thick, more preferablyfrom about 5 to about 1500 μm thick, and even more preferably about 500to about 800 μm thick.

Alternatively, a very thin coating or hydrogel may be applied to a graftor other structure. The coating is applied, e.g., by spraying, so thatthe coating or hydrogel is present on the fibers of the graft. Theinterstices between the fibers are not coated by this technique. Forexample, a wire mesh stent may be coated with a hydrogel without thehydrogel filling the interstices of the mesh. Such a hydrogel or coatingis preferably less than about 10 μm in thickness.

FIG. 7 depicts vascular graft 10 being made by molding hydrogel 14around mandrel 20 and within outer mold 22 while entrapping syntheticmaterial 12 in annulus 26 created by mandrel 20 and outer mold 22. Themonomers and/or macromers used to make hydrogel 14 are poured intoannulus 26 and then polymerized around synthetic material 12. FIG. 8depicts cross-section 8-8 of FIG. 7; hydrogel 14 is shown in phantomlines.

A tubular member of the invention may also be a plastic tubularextrusion of a polymeric material. Such materials could be, for example,silicone polymers, polyester, polyurethane, polypropylene, siliconeelastomer, polytetrafluoroethylene (PTFE) or other suitable materials.The tubular member may also be porous or have a defined permeability,e.g., as a hollow tube fiber with a defined molecular weight cut-off.Porosity/permeability may be incorporated into a tube by the selectionof the material or by adding pores or holes, e.g., by lasers, andpunctures.

Another embodiment of the invention is a method of making a vasculargraft. A first polysaccharide hydrogel tube is made and introducedinside of a tubular member that has an exterior side and an inner sidethat faces the lumen of the tubular member. A second polysaccharidehydrogel tube is made and introduced around the outside of the tubularmember to form a sandwich of a tubular member between two hydrogeltubes. The two hydrogel tubes are then treated to form one unit. Onesuitable treatment is to swell the hydrogel tubes in a solvent to bringthem in to contact with each other and then to chemically react them.Examples of suitable chemical reactions include polymerization,polymeric cross-linking with ultraviolet, heat, or sonic initiators, andchemical cross-linking with gluteraldehyde, or diisocyantes. Thecross-linking agents may be present in the tubes prior to swelling ormay be introduced with the solvent. Suitable solvents include aqueoussolvents, organic solvents, and low boiling point and/or low dielectricconstant organic solvents. One option for forming one unit of thehydrogels is to use a tubular member that is shorter than the twohydrogel tubes so that the two hydrogel tubes are joined around thetubular member. Another option for forming one unit of the hydrogels isto use a tubular member that is porous so that the hydrogels are forcedinto the pores during swelling so that the two hydrogel tubes contacteach other through the pores. As a result, the hydrogels may becomecross-linked through the pores.

The polymers and hydrogels of the invention, whether encapsulating orcoating the tubular member, may if desired, incorporate and slowlyrelease growth factors, thrombolytic drugs, thrombotic drugs, enzymes,restenosis-preventing drugs, inhibitors and other agents used to treatdiseased tissues. Further, gene therapy delivery may be performed bycomplexing gene therapy victors to the polymer or hydrogel, e.g., bycomplexing DNA and a polysaccharide together with a positively chargedion or polymer. Other functions, uses, applications, formulations, andtechnologies of hydrogels known those skilled in the art may be used tocreate further alternative embodiments of the invention.

An embodiment of the invention is a method of making a heparin macromerby reacting heparin with a quaternary ammonium salt to form aheparin-quaternary ammonium salt complex; dissolving theheparin-quaternary ammonium salt complex in an organic solvent with adielectric constant less than the dielectric constant ofdimethylsulfoxide; and decorating the heparin in the heparin-quaternaryammonium salt complex with a polymerizable monomer. Further, a vacuumremoval step of removing the organic solvent with a vacuum may be used,preferably at room temperature.

An embodiment of the invention is a method of making a heparin macromerby reacting heparin with a quaternary ammonium salt to form aheparin-quaternary ammonium salt complex; dissolving theheparin-quaternary ammonium salt complex in an organic solvent that hasa boiling point of less than 190 degree Centigrade at atmosphericpressure; and decorating the heparin in the heparin-quaternary ammoniumsalt complex with a polymerizable monomer. The step of dissolving theheparin-quaternary ammonium salt complex in an organic solvent isalternatively performed with an organic solvent with a boiling point ofless than 114 degrees Centigrade at atmospheric pressure.

The invention includes an optional additional step of decomplexing theheparin quaternary ammonium salt from the heparin-quaternary ammoniumsalt complex by mixing the heparin-quaternary ammonium salt complex witha salt that is not a quaternary ammonium salt

An embodiment of the invention is a method of making a heparin polymerby reacting heparin with a quaternary ammonium salt to form aheparin-quaternary ammonium salt complex; dissolving theheparin-quaternary ammonium salt complex in an organic solvent that hasa boiling point of less than 190 degree Centigrade at atmosphericpressure; decorating the heparin in the heparin-quaternary ammonium saltcomplex with a polymerizable monomer; and polymerizing the monomer tomake a polymer.

An embodiment of the invention is a method of making a heparin hydrogelby reacting heparin with a quaternary ammonium salt to form aheparin-quaternary ammonium salt complex; dissolving theheparin-quaternary ammonium salt complex in an organic solvent with adielectric constant less than the dielectric constant ofdimethylsulfoxide; decorating the heparin in the heparin-quaternaryammonium salt complex with a polymerizable monomer to make a heparinmacromer; and polymerizing the heparin macromer to form a polymer andcross-linking the polymers to form a hydrogel.

An embodiment of the invention is a method of making a polysaccharidemacromer, the method comprising: reacting a polysaccharide with aquaternary ammonium salt to form a polysaccharide-quaternary ammoniumsalt complex; dissolving the polysaccharide-quaternary ammonium saltcomplex in an organic solvent that has a boiling point of less than 190degree Centigrade at atmospheric pressure; and decorating thepolysaccharide in the polysaccharide-quaternary ammonium salt complexwith a polymerizable monomer. The step of dissolving thepolysaccharide-quaternary ammonium salt complex in an organic solventmay also be performed with an organic solvent with a boiling point ofless than 114 degrees Centigrade at atmospheric pressure.

An embodiment of the invention is a method of making a material from apolysaccharide by reacting a polysaccharide with a quaternary ammoniumsalt to form a polysaccharide-quaternary arnmonium salt complex;dissolving the polysaccharide-quaternary ammonium salt complex in anorganic solvent that has a boiling point of less than 190 degreeCentigrade at atmospheric pressure; and decorating the polysaccharide inthe polysaccharide-quaternary ammonium salt complex with a polymerizablemonomer to make a polysaccharide macromer; decorating the polysaccharidein the polysaccharide-quaternary ammonium salt complex with apolymerizable monomer to make a polysaccharide macromer. Moreover theremay be included a step of polymerizing the polysaccharide macromer tomake a polysaccharide polymer. There may further be included a step ofpolymerizing the polysaccharide macromer to form a hydrogel.

An embodiment of the invention is a material for use in a medicalcontext, the material including a hydrogel made of a material includingpolymers, the polymers including heparin polymers made of polymerizableheparin macromers, the hydrogel having covalently cross-linked polymerssuch that the hydrogel remains intact in water. The heparin macromersmay be macromers that are polymerizable while in a solution or in asuspension. The hydrogel may include polymerizable heparin macromerspolymerizable in aqueous solvent and/or polymerizable heparin macromersare polymerizable in organic solvent. The amount of heparin in theheparin hydrogel may be least 1% as measured by dividing the dry weightof heparin macromers by the total dry weight of the hydrogel. And theamount of water in the heparin hydrogel may be at least 5% as measuredby dividing the weight of water in the hydrogel by the total weight ofthe hydrated hydrogel and is preferably in the range of 10%-90% and morepreferably 60%-80% water.

An embodiment of the invention is a material for use in a medicalcontext, the material including a hydrogel made of a material includingpolymers, the polymers including polysaccharide polymers made ofpolymerizable polysaccharide macromers, the hydrogel having covalentlycross-linked polymers such that the hydrogel remains intact in water,and the polysaccharide macromers being macromers that are polymerizablewhile in a solution or in a suspension.

An embodiment of the invention is a material for use in a medicalcontext, the material including a hydrogel including polymers, thepolymers including heparin polymers made of polymerizable heparinmacromers, the hydrogel having covalently cross-linked polymers suchthat the hydrogel remains intact in water, the heparin macromers beingheparin molecules decorated with a monomer chosen from the groupconsisting of monomers polymerizable by free-radical polymerization,monomers polymerizable by addition polymerization, and monomerspolymerizable by condensation polymerization.

An embodiment of the invention is a material for use in a medicalcontext, the material including a hydrogel including polymers, thepolymers including polymers made of polysaccharide macromers, thehydrogel being covalently cross-linked such that the hydrogel remainsintact in water, the polysaccharide macromers being polysaccharidesdecorated with a monomer chosen from the group consisting of monomerspolymerizable by free-radical polymerization, monomers polymerizable byaddition polymerization, and monomers polymerizable by condensationpolymerization.

An embodiment of the invention is a vessel for use in a medical context,the vessel comprising: a tubular member with an inside wall defined byan inner diameter and an outside wall defined by an outside diameterjoined by a thickness, a portion of the cylinder having its inner walland outer wall covered with a hydrogel, the hydrogel including polymers,the polymers including heparin polymers made of polymerizable heparinmacromers, with the hydrogel having covalently cross-linked polymerssuch that the hydrogel remains intact in water.

The hydrogel for the small diameter vascular graft and other suchvessels is preferably at least several μm thick, and more preferably isat least about 50 μm thick, and even more preferably about 500-800 μmthick. The vessel's polysaccharide macromers may be polysaccharidemolecules decorated with a monomer chosen from the group consisting ofmonomers polymerizable by free-radical polymerization, monomerspolymerizable by addition polymerization, and monomers polymerizable bycondensation polymerization. In contrast to certain prior artinventions, the polysaccharide macromers of the present invention may bemacromers that are polymerizable while in a solution or in a suspension.

An embodiment of the invention is a vessel wherein the diameter of theminimum cross-sectional area available for blood flow through the vesselafter the vessel is covered with hydrogel is less than approximately 6.0mm. The vessel may be a fabric vessel that has pores and the hydrogel iscontinuous through a portion of the pores of the fabric vessel.

An embodiment of the invention is a coated stent. Another embodiment isa tissue engineering matrix. Another embodiment is a medical devicecovered to make a biomaterial covering around the device.

The medical device may be coated on one surface, or a portion thereof.Alternatively, the device may be completed coated on all exteriorsurfaces. An encapsulated device is coated on all surfaces with anessentially continuous material. If the material is a hydrogel, thehydrogel is preferably crosslinked so as to have an increased strength.A continuous hydrogel is distinguished from a thin coating because thecoating is attached or adsorbed to the coated surface and is stable solong as that attachment is maintained. But a hydrogel forms a coherentstructure that has stability independent of its attachment to theencapsulated surface. The term encapsulated, as used herein, means tocover. In the case of an encapsulated medical device, the covering isessentially total. In the case of a hydrogel encapsulating a tube, theinside and outside of the tube are covered. The ends of the tube are notnecessarily covered.

An embodiment of the invention is a polymer (or macropolymer) ofvinylpyrrolidone. Vinylpyrrolidone polymers or oligomers may bedecorated with a polymerizable group using techniques known to those ofordinary skill in these arts to make a vinylpyrrolidone macromer that ispolymerizable. The vinylpyrrolidone macromer may be polymerized to makea three-dimensional cross-linked structure. Or the vinylpyrrolidonemacromer may be polymerized to make a larger polymer. Further,cross-links may be incorporated into the larger polymer.

This larger polymer has different properties, e.g., molecular weight,branched structure, cross-linked structure, than the unpolymerizedstarting material. These properties may be manipulated to achieve apolymeric vinylpyrrolidone that has superior adsorbtive properties. Anunbranched, uncrosslinked polymer of vinylpyrrolidone is poorlyadsorbtive. A multi-armed crosslinked or branched polymer, however, ishighly adsorbtive and stable. But a polymer that is too highlycrosslinked will become insoluble and will fall out of solution so thatit is poorly adsorptive and difficult to use in coating techniques.

Techniques for readily determining the molecular weight of largecrosslinked polymers are not available. Therefore it is usuallynecessary to perform routine optimization procedures to developmulti-armed polymers with appropriate branching or crosslink densities.For example, a 100,000 molecular weight vinylpyrrolidone that has beendecorated with between two to ten polymerizable groups is placed intosolution in five samples that vary in concentration by a factor of ten.Each sample is exposed to initiating conditions to react all of thepolymerizable groups. A portion of the surface that is to be coated,e.g., a wire or tube, is exposed to the samples for a set time,preferably between about 2-10 hours. The portion is removed, rinsed inaqueous solution, and tested for adsorption. The sample that caused thehighest amount of adsorption is identified and the procedure is repeatedwith a new concentration range built around the optimal solution.Samples wherein the polymers fall out of solution are rejected as havingpolymers that are too highly crosslinked. A range of parameters may bevaried to ascertain the optimal conditions, including starting polymericmolecular weight, number of polymerizable groups per polymer, andsolution concentrations. This optimization procedure is applicable formaking polymeric, linear, high-molecular weight, and multi-armedpolysaccharides as well as linear, multi-armed, and high molecularweight vinylpyrrolidones.

The vinylpyrrolidone polymer (or macropolymer) is useful for makinglubricious coatings. The coatings may be made on essentially any object.The coatings may be made, for example, by drying a solution ofvinylpyrrolidone polymer, multi-armed polymer, or macromer onto anobject. The macromer may subsequently be polymerized. Alternatively, theobject may be covered with the macromer and the macromer polymerized tomake a coating or encapsulating membrane. The vinylpyrrolidonemacropolymer or the macromer may be combined with othermonomers/polymer/macromers, especially those described herein.

The preferred objects for coating are medical devices or componentsthereof. For example, springs, wires, guide wires, pacemaker leads,stents, implants, antennae, sensors, glucose sensors, tubing, bloodbypass tubing, syringes, catheters, i.v. bags, needles, oxygen tubing,ventricular assist device components, and trochars.

The present invention also has inventive embodiments related to themaking of the materials described herein. One aspect of makingpolysaccharide macromers and the like involves using organicsolvent-soluble mucopolysaccharides. An aspect of this invention is thata complex of the mucopolysaccharide is formed with a cationic moiety,where the complex is organic soluble. The mucopolysaccharide part of theorganic soluble complex can subsequently undergo a variety of chemicalreactions. After completion of the chemical reaction, themucopolysaccharide is de-complexed, producing a chemically modifiedmucopolysaccharide modified in organic solvent (O-MPSAC) that stillretains its functional characteristics. As in the case of heparin, thechemically modified heparin is produced in its active anti-thrombogenicform. The macromers may be used to form the hydrogels or the linear ormulti-armed polymers of the invention.

An aspect of this invention is that a mucopolysaccharide is chemicallymodified in water or dimethylsulfoxide (DMSO) or another equivalentsolvent or solvent mixtures, herein referred to W-MPSAC. The W-MPSAC isthen complexed with a cationic moiety to form a complex that is organicsoluble, which can undergo further chemical reactions. The W-MPSAC maythen be de-complexed.

The schemes shown in FIGS. 9 and 10, scheme I and scheme II,respectively, show these aspects of the invention. The chemical reactionin FIG. 9 can consist of chemically attaching the O-MPSAC to the surfaceof a material that contains reactive species. The chemical reaction inFIG. 9 can also consist of the incorporation of a functional group tothe O-MPSAC that can undergo a further chemical reaction by free radicalprocess or by photo initiated reaction or chemical coupling reaction tovarious polymers.

The first chemical reaction in FIG. 10 can consist of the incorporationof a functional group to the mucopolysaccharide that can undergo afurther chemical reaction by free radical process or by photo initiatedreaction. The first chemical reaction in FIG. 10 can also consist ofchemical coupling reactions to various other chemical moieties.

The second chemical reaction in FIG. 10 can consist of chemicallyattaching the W-MPSAC to the surface of a material that containsreactive species. The second chemical reaction in FIG. 10 can alsoconsist of the incorporation of a functional group to the W-MPSAC thatcan undergo a further chemical reaction by free radical process or byphoto initiated reaction or chemical coupling reaction to variouspolymers.

The use of synthetic materials has gained wide-ranging acceptance inrecent years as suitable materials for medical devices. The extent oftheir application has extended from simple disposable devices, likesyringes, blood bags, catheters, also products like extracorporealdevices, artificial blood vessels, stents, stent grafts to complexartificial organs, such as kidneys, lung, liver, heart assist devicesand implantable devices. These medical devices are required to have theappropriate functional properties, durability, and biological safety.

There is now emerging an additional requirement for these medicaldevices, especially implantable devices, to have biocompatibility withthe biological environment, with minimum or no tissue rejection orreaction. Anti-thrombogenicity is a biocompatibility property that isimportant in many cases.

Prior art for imparting biocompatible properties to medical devicesconsisted of two main routes: (1) certain schemes for attaching amucopolysaccharide (e.g., heparin) to a surface (2) chemicalmodification to introduce groups onto the mucopolysaccharides or thesurface to make them hydrophilic, zwitterionic and/or charged, (e.g.,anionic or cationic).

Route (1) has been achieved by: (I) blending or attaching anorganic-soluble polymer to heparin so that the heparin goes intosolution in organic solvent; (II) treating heparins to dissolve theminto an organic solvent and using the organic solution to coat a medicaldevice, (III) electrostatic binding of heparin to a surface, and (IV)chemically linking the heparin to the surface.

To cite a few examples of these methods, Pusineri et al disclose in U.S.Pat. No. 4,469,827 polymer compositions containing quarterinised aminogroups that ionically bind to heparin. Hsu discloses ionic heparincoating in U.S. Pat. No. 5,047,020, where alkylbenzl ammonium cationsare used to complex with heparin. U.S. Pat. No. 5,541,167, describes acoating composition consisting of stearyldimetylbenzyl ammonium heparincomplex with antifoaming agents. Hsu et al in U.S. Pat. No. 5,417,969inform processes for coating the surface of polyvinylchloride with anorganic solvent soluble solution of heparin complexed with an organiccation. EP 0 769 503 A2, patent application discloses a heparin complexcoating that contains stable ionic bonding and where reduction inanticoagulant activity is minimized. The preferred quaternary ion isalkyldimethylammonium. Yokota and et al have described (in EP 0 781566), an organic soluble heparin complex coating for medical devices,where the cation consists of a quaternary phosphonium moiety. In anotherdisclosure, U.S. Pat. No. 5,270,046, a monomer is formed containingquaternary ammonium groups, which are complexed with heparin. Themonomeric complexed heparin is polymerized with other monomers. Ingeneral, however, these coatings fail under prolonged use inphysiological conditions. The reason for failure is that ions generallydecomplex from the heparin, causing the release of the heparin and thecation.

A preferred embodiment of the present invention is to chemically attachan O-MPSAC to the surface of a medical device that contains reactivespecies. The surface reactive species are able to react with themucopolysaccharide, which is followed by decomplexation, leaving themucopolysaccharide bound to the surface. This is exemplified by formingan organic solvent soluble heparin complex, which can be reacted withbut not limited to, isocyanate or epoxide groups, which have beenincorporated to the surface of the medical device. The isocyanate orepoxide groups are able to react with free hydroxyl of the heparin toform urethane or ether linkage. In another manner the isocyanate orepoxide groups are able to react with the free amino groups of theheparin to form urea or substituted amine linkage. On completion of thesurface reaction the heparin is decomplexed with salt solution, leavingthe heparin chemically bound to the surface in its active form. Theorganic soluble heparin complex can be applied to the medical device by,e.g., dip coating, spray coating or any other coating process.

Another preferred embodiment of this invention is to chemically modifythe mucopolysaccharide component of the organic solvent soluble complex(see FIG. 9), to produce chemically activated O-MPSAC, which isactivated to be able to undergo further chemical reactions. This isexemplified by forming an organic solvent soluble heparin complex, whichcan be reacted with but not limited to, isocyanatoethyl methacrylate ormethacryloyl chloride. The isocyanatoethyl methacrylate or methacryloylchloride is able to react with free hydroxyl of the heparin to formmethacrylate urethane or methacrylate ester linkage with the heparin. Inanother manner the isocyanatoethyl methacrylate or methacryloyl chlorideis able to react with the free amino groups of the heparin to formmethacrylate urea or methacrylate amide linkage with the heparin. Otherpolymerizable groups may be substituted for methacrylates.

The complexed methacrylate heparin of the form described above canundergo free radical polymerization with other macromers and/ormonomers, either in solid state, gel state, in solution, in emulsion, orin suspension. The final polymer, where the heparin complex is attachedto the polymer backbone, can be used to coat medical devices. Theheparin is then decomplexed with salt solution, leaving the activeheparin coated onto the medical device. Other polysaccharides orproteoglycans may be substituted for heparin. And other polymerizablegroups may be substituted for methacrylates.

Polymers made of polysaccharide macromers are distinct from themacromers. The polymers are made of more than one macromer. The polymershave a synthetic polymeric backbone formed by the polymerization of thepolymerizable groups. The polymers have a higher molecular weight thanthe macromers, a factor that affects their solubility and theirstability when adsorbed or otherwise attached to a surface. The polymersmay have a linear, branched, or cross-linked structure. Such polymersmay be distinguished from hydrogels in that a hydrogel is a solid objectthat is incapable of being dissolved or forming a suspension unless itis destroyed, e.g., by grinding. “Synthetic” as used herein means notnaturally found in nature and does not refer to the process whereby anobject is made.

The multi-armed and high molecular weight polymers have superioradsorptive properties compared to unmodified polymers. The optimaldegree of branching or cross-linking that is appropriate is determinableby routine optimization processes as described for preparing lubricioussurfaces. The modified polymers are placed into solution and exposed tothe surface to be modified. The coating may be dried onto the coatedsurface.

Subsequent processing of the coatings of the invention may be used tofurther secure the coating to the surface. The polymers of the coatingmay be prepared to have polymerizable groups or functional groups thatare reactive with the surface. For example, methacrylates,photopolymerizable groups, or isocyanates may be used. Alternatively,the surface may be made chemically reactive, for example by usingepoxide. The chemical groups make bonds with the surface to furthersecure it. Polymerizable groups may be polymerized or reacted with thesurface or with the other crosslinked polymers to further secure thecoating on the surface.

Alternatively, the heparin macromer may be polymerized to form athree-dimensional crosslinked hydrogel. The cross linking may beaccomplished by providing an average of at least two polymerizablegroups per macromer or by mixing the macromers with crosslinkers thathave at least two polymerizable groups.

A complexed methacrylated heparin macromer may be decomplexed with saltsolution, giving a methacrylated heparin macromer that can undergopolymerization with other macromers and/or monomers, by e.g., solution,emulsion, or suspension polymerization. The final polymer, where theactive heparin is attached to the polymer backbone, can be used to coatmedical devices. Thus the macromers may be polymerized to form largermacromers or polymers. Or they may be polymerized to form athree-dimensional hydrogel.

Another embodiment of this invention is a chemically modified O-MPSACwhich is able to undergo photochemical reactions. This is exemplified byforming an organic solvent soluble heparin complex, which can bechemically modified to contain photochemical reactive groups that areable to link to the surface of medical devices. The photochemicalreactive groups can consists of allyl, vinyl, acrylates, methacrylates,azides, nitrenes, carbenes and excited states of ketones, diazo, azocompounds and peroxy compounds, and such groups as are cited in andthose cited in WO 90/00887, which is hereby incorporated herein byreference. An advantage of this method is that photochemical reactivegroups may be bound to the heparin complex in organic solvents, and theresultant product may then be dissolved in organic solvent and coatedonto the medical device. Then applying the appropriate electromagneticradiation to carry out the photochemical reaction; this is followed bydecomplexation, leaving the active heparin bound to the surface.

An advantage of the modified polymers and the polysaccharide macromersof the invention is that they may be reacted with a surface while insolution. For example, modified polymers with polymerizable groups maybe polymerized while in solution. Or photoactivatable groups orelectrophilic groups may be activated while the polymer or macromer isin solution. These approaches are particularly effective when organicsolvents are used because many chemical reactions are much moreefficient in organic solvents as compared to aqueous solvents, e.g.,electrophile-nucleophile reactions. The present disclosure providesnumerous techniques for bringing polysaccharides and modified polymersinto solution so that an effective reaction with the surface may beperformed.

Another preferred embodiment of the invention comprises of the chemicalmodification of the mucopolysaccharide in water or dimethyl sulphoxideor another equivalent solvent or solvent mixtures. This is exemplifiedby the reaction of the activated imidazole carbonate ofpolyethyleneglycol methacrylate (see WO 97/41164, hereby incorporated byreference herein) with heparin to form the heparin-polyethyleneglycolmethacrylate. The heparin-polyethyleneglycol methacrylate is thencomplexed (W-MPSAC, Scheme 2). The W-MPSAC can then undergo furtherchemical reaction in an organic solvent, e.g., by free radical processor by photo initiated reaction. Alternatively, theheparin-polyethyleneglycol methacrylate can undergo chemical reactionsby free radical process or by photo initiated reaction without thecomplexation step. The resultant product may then be complexed to forman organic soluble heparin complex.

EXAMPLE 1 Heparin Complex

5 g sodium heparin (Celsus Laboratories, Inc, USP lyophilized fromporcine intestinal mucosa) was dissolved in 80 ml de-ionized water andallowed to stir for 1 hour in a 250 ml beaker.

8 g benzalkonium chloride (Aldrich Chemical Company, Inc) was allowed todissolve in 80 ml de-ionized water with gentle warming (40-50° C.) on amagnetic stirrer hotplate for 1 hour and then allowed to reach roomtemperature.

To the above vigorously stirred solution of sodium heparin, thebenzalkonium chloride solution was added. A white precipitateimmediately formed and the suspension was further stirred for 1 minute.The precipitate was filtered through a Whatman qualitative filter paper(grade 1).

The white precipitate was collected from the filter paper andre-suspended in 400 ml de-ionized water and allowed to stir for 20minutes. The suspension was filtered as above and re-suspended in 400 mlde-ionized water and filtered again. The precipitate was once againsuspended in 400 ml de-ionized water and then poured into a dialysismembrane Celli Sep:MWCO 3,500 and dialyzed against 10 L de-ionized waterfor a minimum of 16 hours.

The precipitate was collected and dried on a glass dish in a vacuum ovenat 60° C. for 12 hours.

Dry gray-yellow crystals were obtained with a yield of 10 g.

EXAMPLE 2

3 g 4,4′-methylenebis(phenyl isocyanate) (MDI) (Aldrich ChemicalCompany, Inc) was dissolved in 100 ml anhydrous tetrahydrofuran (THF).Polyurethane tubing was dipped in to the above solution for 30 secondsand was then allowed to stir-dry at 60° C. for 1 hour.

5 g of dry crystals of complexed heparin from Example 1 were dissolvedin 100 ml anhydrous dichloromethane (DCM). The MDI coated polyurethanetubing was dipped into the DCM solution of complexed heparin for 30seconds and then allowed to air-dry for 2 hours.

The tubing was dyed with 0.075% w/v pH 8.5 toludine blue aqueoussolution for 30 seconds and washed with de-ionized water. A very faintpurple color due to the complexed heparin was observed. The tubing wasthen immersed in 25% w/v solution of sodium chloride at 40° C. for 30minutes. The tubing was washed with de-ionized water and again dyed. Anintense dark purple color due to complexed heparin was observed.

In a similar experiment where the polyurethane tubing was not initiallycoated with MDI and was de-complexed in sodium chloride solution (asabove) no purple coloration due to the heparin was observed.

EXAMPLE 3

1. Heparin Methacrylate (methacryloyl chloride)

5 g of dry crystals of complexed heparin from Example 1 were dissolvedin 100 ml anhydrous DCM in a 250 ml quickfit conical flask. To this wasadded 0.1265 g (1.25×10⁻³ moles) triethylamine.

Methacryloyl chloride (Aldrich Chemical Company, Inc) was distilledunder reduced pressure to obtain a very pure sample. 0.1306 g (1.25×10⁻³moles) of the above distilled methacryloyl chloride was dissolved in 30ml anhydrous DCM and placed in a stoppered quickfit pressure equalizingfunnel above the vigorously stirred solution of DCM containing thecomplexed heparin. The methacryloyl chloride solution was addeddrop-wise over a period of 30 minutes to the complexed heparin solutionin DCM.

The DCM was rotary evaporated and the complexed heparin methacrylate wasdried in a vacuum oven at 40° C. for 2 hours.

Complexed methacrylate was characterized by ¹H and ¹³C.

0.2 g of the above complexed heparin methacrylate was dissolved in 10 ml2-hydroxyethylmethacrylate and to this was added 0.02 g ethylene gluoldimethacrylate and 0.02 g 2,2′-azobis(2,4-dimethylvaleronitrile)(Dupont). The above clear solution was degassed for 30 minutes.

The above polymerization mixture was poured into a polypropylene concavemold and then a polypropylene convex mold was placed onto the concavemold allowing the excess solution to overflow, thereby uniformly fillingthe space between the concave and convex molds. The sealed molds werethen heated to a temperature of 65° C. for 4 hours and then at 110° C.for 1 hour.

The molds were cooled and opened to obtain a clear dehydrated rigidhydrogel. These were then hydrated in de-ionized water for 10 hours,after which they were placed in boiling solution of 25% w/v sodiumchloride aqueous solution for 1 hour and then equilibrated in de-ionizedwater. The hydrogel was dyed (as in Example 2) and a uniform intensedark purple coloration was observed throughout the hydrogel.

In a similar experiment where no methacrylate of the heparin was formed,a hydrogel of the complexed heparin alone was formed. After boiling in25% w/v sodium chloride aqueous solution and equilibrating in water thehydrogel was dyed. Dark purple precipitated particles of heparin couldbe observed on the surface of the hydrogel and were easily washed awaywith de-ionized water, leaving a blue coloration to the hydrogel. Thiscoloration is identical to the hydrogels formed without any complexedheparin.

2. Heparin Methacrylate (isocyanatoethylmethacrylate)

5 g of dry crystals of complexed heparin from Example 1 were dissolvedin 100 ml anhydrous DCM in a 250 ml thick-walled glass bottle with cap.To this was added 0.194 g (1.25×10⁻³ moles)2-isocyanatoethylmethacrylate (Aldrich Chemical Company, Inc) and 0.05 gdibutyltin dilaurate (Aldrich Chemical Company, Inc). The cap wasscrewed on tight and the solution was stiffed for 16 hours at 40° C.

The DCM was rotary-evaporated off and the product dried in a vacuum at40° C. for 2 hours.

As in Example 3(1), hydrogels were made and the heparin was decomplexedand dyed. Again, a uniform intense dark purple coloration was observedthroughout the hydrogel whereas the complexed heparin with nomethacrylate coupling heparin particles precipitated on the surface ofthe hydrogel and were easily washed away with de-ionized water.

EXAMPLE 4

5 g of Heparin methacrylate from example 3 i) was dissolved in 100 ml of2-propanol. To this was added 20 g methoxypolyethlyeneglycol 2000methacrylate (MPEG 2000 MA) (Inspec U.K.) and 3 g of2-hydroxyethylmethacrylate.

A 250 ml, 3-necked reaction vessel equipped with stirrer, thermometer,condenser and nitrogen inlet tube was charged with 100 ml 2-propanol. Tothe 2-propanol was added 5 g of heparin methacrylate [from example 3(i)], 20 g methoxypolyethyleneglycol 2000 methacrylate (MPEG 2000 MA)(Inspec U.K.) and 3 g 2-hydroxyethylmethoxylate. (HEMA) The 250 ml, 3necked reaction vessel was placed in a silicone oil bath at 120° C. andthe 2-propanol was stirred gently and nitrogen was bubbled through thesolution (100 cm³/min).

When the temperature in the 250 ml, 3-necked reaction vessel reached 75°C., 0.25 g 2,-2′-azobis(2,4-dimethylvaleronitrile) was added and thestirrer speed was increased to 750 rpm. After approximately 15 minutes avery viscous solution was obtained and the reaction was allowed tocontinue for 30 minutes, periodically added 30 ml 2-proponal to diluteand reduce the viscosity of the solution. In total four 30 ml aliquotsof 2-propanol was added.

The polymer was cooled down to room temperature and the isopropanolrotary evaporated off and the polymer dried in a vacuum oven at 40° C.for 8 hours.

GPC (Gel permeation chromatography) showed that the average molecularweight of the polymer was approximately 400,000 when using polyethyleneglycols as standards.

EXAMPLE 5

3 g MDI was dissolved in 100 ml of anhydrous THF. A 150 mm long and 2 mmdiameter polyurethane tube was dipped into the above solution for 30seconds and was then allowed to air dry at 60° C. for 1 hour.

5 g of heparin copolymer from Example 4 was dissolved in anhydrous DCM.The MDI coated polyurethane tubing was dipped into the DCM solution ofcopolymer for 30 seconds and was then allowed to air at 60° C. for 30minutes and then at room temperature for 16 hours.

The tubing was immersed in 25% w/v sodium chloride solution at 40° C.for 30 minutes. The tubing was washed with de-ionized water and driedwith toludine blue (example 2). A homogenous intense dark purplecoloration was obtained on the tubing. In addition the polyurethanetubing was completely wetted and very lubricious. The lubricity did notdiminish even when it was rubbed between forefingers and thumb 20 times.The tubing was then immersed in phosphate buffered saline at 50° C. for16 hrs and then rubbed between the forefinger and thumb. Again there wasno observable discharges in lubricity.

EXAMPLE 6

As in Example 4, 5 g heparin methacrylate was copolymerized, except thatinstead of MPEG 2000 MA and HEMA, 20 g N-vinylpyrrolidone was used asthe co-monomer in 2-propanol. The polymerization was allowed to continuefor 1 hour.

The average molecular weight of the copolymer was approximately 300,000as determined by GPC.

As in Example 5 a 150 mm long and 2 mm diameter polyurethane tubing wasdipped into a 3% w/v MDI in THF solution and air dried at 60° C. for 1hour.

The tubing was then dipped in a 5% w/v heparin-vinylpyrrolidonecopolymer in DCM for 30 seconds and then air dried at 60° C. for 2hours.

The tubing was then immersed in 25% w/v sodium chloride solution at 40°C. for 30 minutes. The tubing was washed with de-ionized water and dyedwith toludine blue. A homogenous intense dark purple colorization wasobtained. The polyurethane tubing was completely wetted and extremelylubricious. The lubricity was equivalent to that obtained for in Example5.

EXAMPLE 7

5 g heparin methacrylate was copolymerized with 20 g acrylic acid in2-proponal. The reaction conditions were the same as in Example 6.

The average molecular weight was determined to be 400,000 using GPC andpolyacrylic acids as standards.

As in Example 5 a 150 mm long and 2 mm diameter polyurethane tubing wasdipped into a 3% w/v MDI in THF solution and air dried at 60° C. for 1hour.

The tubing was dipped into a 5% w/v heparin-acrylic acid copolymer inmethanol/dimethylacetamide solution (90:10) for 30 seconds and the airdried at 60° C. for 3 hours.

The lubricity of the surface of the polyurethane tubing was very similarto that in Example 5 and 6 after washing in 25% w/v sodium chloridesolution (as performed in Example 5 and 6).

XPS confirmed the presence of heparin on the surface as SO₄ groups bedetached on the surface of the polyurethane tubing.

EXAMPLE 8

PVC tubing 150 mm long 2 mm in diameter were coated with the heparincopolymers synthesized in Example 4, 6 and 7. The coating conditions forthe respective heparin copolymers were identical to the Examples 5, 6and 7 respectively.

In all cases a very durable lubricious coating (when wet) was obtainedand the presence of heparin on the surface of the PVC was detectedeither using toludine blue or XPS.

EXAMPLE 9

A polyurethane tube (150 mm long, 2 mm diameter) was dipped into a 2%w/v solution of poly [1,4-phenylenediisocyanate-co-poly(1,4-butanediol)]diisocyanate (Aldrich Chemical Co.) in anhydrous THFand allowed to air dry at 60° C. for 2 hours.

2 g of heparin complex (from Example 1), 2 g polyethylene oxide (M. W.100,00) (Aldrich Chemical Co.) and 0.25 g MDI were dissolved in 100 mlanhydrous DCM. The polyurethane tubing was dipped with the abovesolution for 30 seconds and allowed to air dry at 60° C. for 1 hour andthen at 22° C. for 16 hours.

The polyurethane tubing was immersed in 25% w/v sodium chloride solutionat 40° C. for 30 minutes and then washed with de-ionized water and thendyed with toludine blue. An intense homogenous dark purple colordeveloped on the tubing and when wet the tubing was highly lubricious.The lubricity was comparable to in Example 5, 6 and 7.

EXAMPLE 10

A similar experiment to the one in Example 9 was conducted withpolyvinylpyrrolidone (M. W. 1,300,000) (Aldrich Chemical Co.) with allthe experimental conditions the same. Again the results showed thatheparin was present on the surface and the polyurethane tubing washighly lubricious and comparable to the coating in Example 9.

EXAMPLE 11

A PTFE tube measuring 3.8 mm O.D; length 8 cm was placed on a rotatablemandrel. Placed over the PTFE was a knitted fabric tube measuring 4.0 mmI.D; length of 7 cm.

A methacrylated polyvinylpyrrolidone (MPVP) was made by dissolving 6 gof polyvinylpyrrolidone (PVP) in dichloromethane. A solution ofisocyanato ethyl methacrylate (0.8 g dissolved in 20 ml ofdichloromethane) was added drop wise to the stirred solution of PVP. Thereaction was allowed to proceed for a further 2 hours. Dichloromethanewas then rotary evaporated and the MPVP formed was used in the vasculargraft formulation as per the following: Hydoxyethyl Methacrylate 18 g;Heparin Methacrylate (as from example 3) 1 g; MPVP 0.6 g; Ethyleneglycol dimethacrylate 0.08 g; thermal initiator 0.12 g (Vazo 52).

The vascular graft formulation was then added drop wise onto therotating knitted fabric until a homogeneous viscous film was formed,which totally encapsulated the fabric. Then the graft was exposed to UVlight from a medium pressure mercury arch lamp for 10 minutes; then thegraft was placed in a vacuum oven for 3 hours at 70° C.

The heparin contained in the graft was de-complexed with saturated NaClsolution. The final vascular graft had a hydrogel that totallyencapsulated the fabric, which had a water content of 50%. Heparinactivity was measured by antithrombin binding assay and was found to be5 units per 100 mm².

EXAMPLE 12

This example shows that vascular grafts made as described herein areeffective for use with patients. A 28 day animal study was carried outon two pigs. The vascular grafts as made according to Example 11 wereimplanted in the carotid artery of each animal using aseptic surgicaltechniques known to those skilled in the art. The carotid arties werevisualized at day nine, and day 28 using a medical X-ray arteriogram.The arteriogram showed that the grafts were patent with endothelialcells, were clean, and showed no evidence of thrombus at both days.

The embodiments of the invention set forth herein are exemplary and notintended to be limiting in scope. Patents, patent applications, andarticles mentioned in this application are hereby incorporated herein byreference.

1. A method of preparing a medical device coating comprising exposingthe medical device to an aqueous solution comprising a first polymer anda polymeric heat-initiated crosslinker and heating the solution tothereby crosslink the first polymer with the heat-initiated polymericcross-linker and form the coating on the medical device, wherein thefirst polymer comprises heparin pendant groups, nucleophilic pendantgroups, and carboxylic acid pendant groups.
 2. The method of claim 1wherein the first polymer further comprises sulfate groups and/orsulfonate groups.
 3. The method of claim 1 further comprisingpreparation of the first polymer by a reaction comprising polymerizationa heparin macromer.
 4. The method of claim 3 wherein the polymerizationof the heparin macromer further comprises copolymerization with apolyethylene glycol macromer.
 5. The method of claim 3 wherein themedical device is chosen from the group consisting of a wire, a tube, acoil, a spring, stent, catheter, balloon, and an implantable device. 6.The method of claim 3 wherein preparation of the first polymer comprisesreacting heparin with a quaternary ammonium salt to form aheparin-quaternary ammonium salt complex and then dissolving saidcomplex in an organic solvent, followed by reaction of the complex witha monomer in the solvent to form the macromer.
 7. The method of claim 6further comprising, before reaction with the medical device, removingthe quaternary ammonium salt by exposure to another salt.